Implantable grafts are commonly used in treatment of diseased blood vessels. One such device is a synthetic vascular graft designed to replace damaged or dysfunctional tissue. Such damage or dysfunction can arise, for example, from arterial or venous pathways that have been damaged by thrombosis, an aneurysm or occlusion. The graft provides an artificial lumen through which blood may flow.
Natural blood vessels are often damaged during treatment of renal failure. For example, when treating patients with renal failure using dialysis, it is necessary to have ready access to blood vessels in order to continuously withdraw blood from the patient in amounts of over 200 ml/min. For dialysis to be effective, it must be repeated on a regular schedule of two or more treatments per week. Each time, a vein is accessed using a relatively large bore needle. As a result of the repeated percutaneous access, the vein will often collapse along the puncture tract or become aneurismal, leaky, or filled with clot. The latter can cause significant risk of pulmonary embolism. As a result, in dialysis treatments, artificial grafts have been used as an alternative to using a patient's own veins, in an attempt to avoid these complications.
Thus, one increasingly useful application of a vascular prosthesis is as a bypass shunt between an artery and a vein. A graft surgically placed between an artery and a vein (AV fistula) is commonly used in dialysis patients. This bypass or fistula is particularly useful for allowing multiple needle access, as is required for hemodialysis treatments.
Grafts can be made from a variety of materials such as textiles and formed polymers. Vascular grafts are often made from polytetrafluoroethylene (PTFE) tubes, and in particular, from expanded polytetrafluoroethylene (ePTFE) tubes. When PTFE is expanded or stretched to form tubes, the material consists of a unique microstructure of nodes interconnected by small fibrils. The space between the nodes that is spanned by the fibrils is defined as the internodal distance (IND). By varying the conditions of manufacture of the ePTFE tubes, such as temperature and rate of stretching and expansion, it is possible to vary the space between the nodes and the number and diameter of fibrils. Expanded PTFE is particularly suitable as an implantable prosthesis as it exhibits the desirable characteristics of superior biocompatibility and low thrombogenicity.
Expanded PTFE products that are stretched and expanded at high temperatures and rates are more homogeneous in structure. The IND is smaller and there are a greater number of fibrils in the ePTFE tubes. As a result, the product is stronger than if it had been made at lower temperatures and/or slower rates. In addition, the porosity is reduced. By varying the conditions of manufacture, it is possible to often obtain a final product having desired porosity, strength, and flex qualities.
It is a goal in graft technology to mimic, as closely as possible, the natural function of the blood vessel being replaced. This involves finding a graft material and design that will be sufficiently strong to resist tear and other mechanical damage, to be sufficiently flexible and compliant to accommodate the natural variability of flow and pressure of blood, and to be sufficiently porous to allow for enhanced healing and appropriate tissue ingrowth to anchor the prosthesis within the blood vessel and integrate it within the body.
The internal structure of ePTFE is desirable in a number of respects. The diameter of the fibrils formed in ePTFE is much smaller than the diameter of fibers of knitted or woven fabrics that have been used previously in vascular prostheses. Expanded PTFE tubes having a relatively large IND also possesses a higher degree of porosity than PTFE. These characteristics create a better substrate for cellular ingrowth, improved flexibility, and greater compliance in a graft. As a result, a prosthesis formed of ePTFE can more closely approximate the natural function of the blood vessel being replaced. Consequently, reduced thrombogenicity, reduced incidence of intima hyperplasia, and improved cellular ingrowth can be expected from ePTFE grafts as compared to a prosthesis formed of other presently available materials or unexpanded PTFE.
Current graft materials and designs have not fully achieved the desired result of mimicking natural vessels, and disadvantages of using the presently available ePTFE grafts remain. For example, when the IND is large so as to increase porosity and improved ingrowth, then the radial tensile strength of the tube is reduced as is the ability of the tube to retain sutures used during implantation. Such microporous tubes tend to exhibit low axial tear strength, so that a small tear or nick will tend to propagate along the length of the tube. Thus, there is a trade-off between optimal porosity and flexibility, and optimal strength.
In addition to the usual structural limitations of using ePTFE for grafts, there is an additional disadvantage of using implantable ePTFE vascular grafts as access shunts for hemodialysis. Specifically, it is difficult to elicit natural occlusion of suture holes created during implantation. As a result, the PTFE grafts are generally not used to withdraw blood until they have been in place for a minimum of 14 days after surgery. This time is required to allow time for protective ingrowth tissue to form and keep blood from leaking from the suture holes. Use of the graft before this period may result in complications such as a hematoma surrounding the graft, false aneurysm, and possibly graft occlusion. Thus, in order to maintain the integrity of the graft, blood cannot be withdrawn from a PTFE vascular graft until the suture holes have healed. However, waiting this amount of time to treat a dialysis patient causes undesirable build-up of toxins in the blood with its attendant problems.
A further problem associated with grafts used for hemodialysis is that repeatedly piercing the graft can compromise its integrity, causing large-scale tears in some instances, or more often result in hematomas where small amounts of blood leak from the needle entry point. A number of designs for ePTFE vascular grafts have been developed to address these problems.
For example, U.S. Pat. No. 4,619,641 discloses a two-piece coaxial double lumen arteriovenous graft. This graft consists of an outer tube positioned over an inner tube, the space between being filled with a self-sealing adhesive. The self-sealing adhesive helps prevent hematomas caused by piercing the graft. A disadvantage of this design is that completely filling the space between tubes with adhesive limits its flexibility and compliance.
In an attempt to increase radial tensile and axial tear strength of ePTFE tubes, U.S. Pat. No. 4,743,480 discloses a method of altering the extrusion process so as to reorient the fibrils in the node and fibril matrix.
U.S. Pat. No. 6,053,939 discloses a single layer ePTFE graft which releases heparin after grafting. Spaces between the nodes and fibrils are chemically treated to make the inner surface of the tube hydrophilic. Tissue-inducing substances and anti-thrombotic substances (such as heparin) are then covalently bonded to the hydrophilic inner surface of the tube and pores. The result is a high patency ratio and reduced risk of thrombosis. Although increased patency is achieved using this technology, there is still a period of delay before the graft can safely be used for dialysis. In addition, there is still a risk of hematoma caused by repeated piercing of the graft during dialysis.
U.S. Pat. No. 5,192,310 discloses a vascular graft having a primary lumen and at least one secondary lumen which share a common side wall. The secondary lumen is filled with a self-sealing, non-biodegradable, biocompatible polymer. However, this graft is difficult to make using traditional extrusion methods. The graft is made by using unconventional methods, involving a combination extrusion and injection molding process. As a result, the manufacture of this graft is expected to result in a non-uniform and irregular pattern of nodes and fibrils. This irregular conformation becomes problematic during the sintering step during which time melt fractures and other inconsistencies in the microstructure will occur. Thus, this disclosed method of making the graft appears unreliable, costly and likely to produce a defective product.
Thus, there is a need for a graft which provides desirable porosity, resists tears at suture holes, and resists blood flow through puncture holes caused by repeated needle access.